|Publication number||US7904134 B2|
|Application number||US 12/070,521|
|Publication date||Mar 8, 2011|
|Filing date||Feb 19, 2008|
|Priority date||Jul 7, 2004|
|Also published as||EP1769418A1, US7346382, US7680526, US7860548, US8379952, US8983155, US9135400, US9235685, US20060017749, US20080114233, US20080114579, US20080154341, US20090208073, US20090210208, US20110196253, US20130030276, US20150045852, WO2006017053A1|
|Publication number||070521, 12070521, US 7904134 B2, US 7904134B2, US-B2-7904134, US7904134 B2, US7904134B2|
|Inventors||Cameron C. McIntyre, Christopher R. Butson, John D. Hall, Jaimie M. Henderson|
|Original Assignee||The Cleveland Clinic Foundation|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (132), Non-Patent Citations (189), Referenced by (38), Classifications (17), Legal Events (2)|
|External Links: USPTO, USPTO Assignment, Espacenet|
This application is a continuation of U.S. application Ser. No. 10/885,982, filed Jul. 7, 2004, now U.S. Pat. No. 7,346,382 which is hereby incorporated herein by reference in its entirety.
This patent application pertains generally to neurosurgery and more particularly, but not by way of limitation, to brain stimulation models, systems, devices, and methods.
High frequency deep brain stimulation (DBS), such as of the thalamus or basal ganglia, represents a clinical technique for the treatment of disorders such as essential tremor and Parkinson's disease (PD). Pilot studies have also begun to examine the utility of DBS for treating dystonia, epilepsy, and obsessive-compulsive disorder. However, understanding of the therapeutic mechanisms of action remains elusive. It is also unclear what stimulation parameters, electrode geometries, or electrode locations are better suited for existing or future uses of DBS.
A DBS procedure typically involves first obtaining preoperative images of the patient's brain, such as by using a computed tomography (CT) scanner device, a magnetic resonance imaging (MRI) device, or any other imaging modality. This sometimes involves first affixing to the patient's skull spherical or other fiducial markers that are visible on the images produced by the imaging modality. The fiducial markers help register the preoperative images to the actual physical position of the patient in the operating room during the later surgical procedure.
After the preoperative images are acquired by the imaging modality, they are then loaded onto an image-guided surgical (IGS) workstation, such as the StealthStation® from the Surgical Navigation Technologies, Inc. (SNT) subsidiary of Medtronic, Inc., for example. Using the preoperative images being displayed on the IGS workstation, the neurosurgeon can select a target region within the brain, an entry point on the patient's skull, and a desired trajectory between the entry point and the target region. The entry point and trajectory are typically carefully selected to avoid intersecting or otherwise damaging certain nearby critical brain structures.
In the operating room, the patient is immobilized and the patient's actual physical position is registered to the preoperative images displayed on the IGS workstation, such as by using a remotely detectable IGS wand. In one example, the physician marks the entry point on the patient's skull, drills a burr hole at that location, and affixes a trajectory guide device about the burr hole. The trajectory guide device includes a bore that can be aimed using the IGS wand to obtain the desired trajectory to the target region. After aiming, the trajectory guide is locked to preserve the aimed trajectory toward the target region.
After the aimed trajectory has been locked in using the trajectory guide, a microdrive introducer is used to insert the surgical instrument along the trajectory toward the target region of the brain. The surgical instrument may include, among other things, a recording electrode leadwire, for recording intrinsic electrical brain signals, a stimulation electrode leadwire, for providing electrical energy to the target region of the brain, or associated auxiliary guide catheters for steering a primary instrument toward target region of the brain. The recording electrode leadwire is typically used first to confirm, by interpreting the intrinsic electrical brain signals, that a particular location along the trajectory is indeed the desired target region of the brain. The stimulation electrode leadwire, which typically includes multiple closely-spaced electrically independent stimulation electrode contacts, is then introduced to deliver the therapeutic DBS stimulation to the target region of the brain. The stimulation electrode leadwire is then immobilized, such as by using an instrument immobilization device located at the burr hole entry in the patient's skull. The actual DBS therapy is often not initiated until a time period of about two-weeks to one month has elapsed. This is due primarily to the acute reaction of the brain tissue to the introduced DBS stimulation electrode leadwire (e.g., the formation of adjacent scar tissue), and stabilization of the patient's disease symptoms. At that time, a particular one of the stimulation electrode contacts is then selected for delivering the therapeutic DBS stimulation, and other DBS parameters are adjusted to achieve an acceptable level of therapeutic benefit. However, these parameter selections are typically currently achieved via arbitrary trial-and-error, without visual aids of the electrode location in the tissue medium or computational models of the volume of tissue influenced by the stimulation.
The subthalamic nucleus (STN) represents the most common target for DBS technology. Clinically effective STN DBS for PD has typically used electrode contacts in the anterior-dorsal STN. However, STN DBS exhibits a low threshold for certain undesirable side effects, such as tetanic muscle contraction, speech disturbance and ocular deviation. Highly anisotropic fiber tracks are located about the STN. Such nerve tracks exhibit high electrical conductivity in a particular direction. Activation of these tracks has been implicated in many of the DBS side effects. However, there exists a limited understanding of the neural response to DBS. The three-dimensional (3D) tissue medium near the DBS electrode typically includes both inhomogeneous and anisotropic characteristics. Such complexity makes it difficult to predict the particular volume of tissue influenced by DBS.
A treating physician typically would like to tailor the DBS parameters (such as which one of the stimulating electrodes to use, the stimulation pulse amplitude, the stimulation pulse width, or the stimulation frequency) for a particular patient to improve the effectiveness of the DBS therapy. This is a complex problem because there are several different DBS parameters than can be varied. Because selecting a particular DBS electrode contact and parameter combination setting is typically a trial-and-error process, it is difficult and time-consuming and, therefore, expensive. Moreover, it may not necessarily result in the best possible therapy or in avoiding the above-mentioned undesirable side effects. Therefore, there is a need to provide help to speed or otherwise improve this DBS parameter selection process or to otherwise enhance DBS techniques.
The patent or application file contains at least one drawing executed in color. Copies of this patent or patent application publication with color drawing(s) will be provided by the Office upon request and payment of the necessary fee.
In the drawings, which are not necessarily drawn to scale, like numerals describe substantially similar components throughout the several views. Like numerals having different letter suffixes represent different instances of substantially similar components. The drawings illustrate generally, by way of example, but not by way of limitation, various embodiments discussed in the present document.
The following detailed description includes references to the accompanying drawings, which form a part of the detailed description. The drawings show, by way of illustration, specific embodiments in which the invention may be practiced. These embodiments, which are also referred to herein as “examples,” are described in enough detail to enable those skilled in the art to practice the invention. The embodiments may be combined, other embodiments may be utilized, or structural, logical and electrical changes may be made without departing from the scope of the present invention. The following detailed description is, therefore, not to be taken in a limiting sense, and the scope of the present invention is defined by the appended claims and their equivalents.
In this document, the terms “a” or “an” are used, as is common in patent documents, to include one or more than one. In this document, the term “or” is used to refer to a nonexclusive or, unless otherwise indicated. Furthermore, all publications, patents, and patent documents referred to in this document are incorporated by reference herein in their entirety, as though individually incorporated by reference. In the event of inconsistent usages between this document and those documents so incorporated by reference, the usage in the incorporated reference(s) should be considered supplementary to that of this document; for irreconcilable inconsistencies, the usage in this document controls.
1. Modeling Techniques
One fundamental step toward understanding the neural response to DBS is characterizing the electric field generated by a stimulating electrode. The electric field is dependent on the shape of the electrode and the electrical conductivity of the tissue medium. DBS electrodes are three-dimensional structures and the tissue conductivity of the central nervous system (CNS) is both inhomogeneous (dependent on location) and anisotropic (dependent on direction). The tissue inhomogeneity and anisotropy surrounding the DBS electrode can alter the shape of the electric field and the subsequent neural response to stimulation. Therefore, in one example, we employed diffusion tensor imaging (DTI) to estimate the electrical conductivity tensor of the tissue medium surrounding one or more DBS electrodes. We incorporated the tissue conductivity data into a finite element model (FEM) tailored to accurately represent the structure of the particular clinical DBS electrode and surrounding tissue medium. We then used these models to predict the volume of tissue likely to be affected by typical stimulation parameters (e.g., stimulation pulse amplitude of between about 1 and 3 Volts, a stimulation pulsewidth of about 0.1 ms, and a stimulation frequency of about 150 Hz). We refer to this volume of t issue likely to be affected as the “volume of influence” (VOI) or “volume of activation” (VOA).
B. Exemplary Methods
We developed, among other things, three-dimensional finite element models (FEMs) of the Medtronic 3387-89 DBS lead (Medtronic, Inc., Minneapolis, Minn.). We examined at least two representations of the nearby tissue electrical properties. In one example, the finite element model used a constant homogenous isotropic tissue conductivity of about 0.3 S/m. In another example, the finite element model explicitly represented tissue anisotropy with conductivity tensors (σ) derived from diffusion tensor magnetic resonance images. Both examples of finite element models used an about 0.2 mm thick sheath of encapsulation tissue (modeled with a conductivity of about 0.15 S/m) about the DBS electrode leadwire shaft. In one example, the FEM was implemented using 231,404 elements in a commercially available software package such as FEMLAB 2.3 (COMSOL Inc., Burlington, Mass.). In one example, a 100×100×100 mm3 cube about the electrode contact was used as a FEM model boundary, which was set to a boundary condition of 0 V. In this example of the FEM model, the electrode contact was set to a boundary condition of the DBS stimulus voltage. The potential distribution (Ve) generated in the tissue medium was calculated from the Laplace equation:
∇·σ∇V e=0, (Eq. 1)
using a Good Broyden iterative solver and algebraic multigrid preconditioner. Doubling the density of the FEM mesh or doubling the distance of the boundary from the electrode (i.e., quadrupling the size of the 100×100×100 mm3 tissue box) yielded a potential distribution Ve that differed only by less than 2% when compared to the default model.
Diffusion tensor imaging (DTI) characterizes the diffusional behavior of water in tissue on a voxel-by-voxel basis in terms of a matrix quantity from which the diffusion coefficient can be obtained corresponding to any direction in space. The electrical conductivity tensor (σ) of a tissue medium is obtainable from the corresponding diffusion tensor (D). The hypothesized relationship between electrical conductivity and water diffusion in tissue is prompted by the observation that in a structured medium the two processes are related through mutual respect for the boundary conditions imposed by the tissue geometry. In our example, the conductivity tensor σ was directly solved for at each voxel using a linear transform of D:
σ=(σe /d e)D, (Eq. 2)
where σe is the effective extracellular conductivity and de is the effective extracellular diffusivity. Our example used a ratio of ((σe/de)=0.736 (S−s)/mm2) as determined from published experimental and empirical data.
In one example, the DTI data was acquired using a 1.5 T Philips Gyroscan NT using a single-shot echo-planar imaging (EPI) sequence with the SENSE parallel imaging scheme (SENSitivity Encoding, reduction factor R=2.5). In this example, the imaging matrix was 96×96 with a field of view of 240×240 mm, which was zero-filled to 256×256. In this example, axial slices of 2.5 mm thickness were acquired parallel to the anterior-posterior commissure line. In this example, the diffusion weighting was encoded along 30 independent orientations and the b-value was 700 s/mm2. In this example, the dorsal STN was located in axial slices using stereotactic coordinates and co-registration with the Schlatlenbrand and Bailey  brain atlas. We extracted the DTI data from the 10×10 mm region that surrounded our electrode location in the STN (in this example, the electrode was located 2 mm ventral, 10 mm lateral, and 1 mm posterior to the mid-commissural point). We then transformed the diffusion tensors to conductivity tensors, as discussed above. We then incorporated the conductivity tensors into co-registered subdomains in the FEM. Then, using the FEM, we solved for the potential distribution generated in the tissue medium by DBS, as discussed above.
C. Exemplary Results
Using the exemplary methods discussed above, we compared the electric field of a theoretical point source, a DBS electrode in an isotropic medium, and a DBS electrode in an anisotropic medium representative of the STN and surrounding tissue structures.
The top row of
The middle row of
For the point source, Δ2Ve vertical (i.e., the left middle picture of
In the example of
Because Δ2Ve represents the effective volume of activation of nearby tissue, this model can be used to adjust the electrode location or stimulation parameters to obtain a desired volume of activation for the DBS stimulation, such as to activate substantially the entire STN, as illustrated by the location and the −3V stimulation in the bottom far right panel of
During extracellular stimulation of the CNS, axonal elements typically represent the most excitable components of neurons near the electrode. Evaluation of Δ2Ve can provide qualitative predictions on the likelihood of neural activation by an extracellular source. Therefore, to provide a quantitative reference to the Δ2Ve data in
This analysis revealed that for a 150 Hz train of cathodic stimuli 0.1 ms in duration a Δ2Ve>12 mV always generated propagating action potentials at the stimulus frequency. When the axon was further than 3 mm from the electrode, a Δ2Ve>8 mV was enough for activation. Also, in this example, the axon model never blocked firing during −3 V; 0.1 ms; 150 Hz stimulation for any of the positions examined.
D. Discussion of Exemplary Results
DBS represents an effective clinical therapy for movement disorders. However, the existing limited understanding of the effects of stimulation on the underlying neural tissue hinders future advancement of this technology. The electric field generated by one or more DBS electrodes, using therapeutic DBS stimulation parameters, can directly activate a large volume of tissue, as illustrated by
The present model provides quantitative results on the effects of DBS that would be difficult to achieve experimentally. Like most models, however, they involved some simplifying approximations worth noting. First, we used electrostatic analysis and the resolution of our diffusion tensor based tissue conductivities was on the order of 1 mm. In general, however CNS tissue typically has a small reactive component that results in slight increases in conductivity at higher frequencies. Also, micro-inhomogeneities exist on scales smaller than the 1 mm. However, a reactive component or higher resolution diffusion tensor based conductivity could be used with the present model techniques, if desired.
Second, neural activation that results from applied fields could be more accurately predicted by directly coupling the electric field data to multi-compartment cable models of individual neurons. The present model techniques, however, provide easier estimation of the volume of tissue supra-threshold, and our estimation is derived directly from the field data. By evaluating Δ2Ve in a plane containing the electrode contact, one can conceptualize the spatial characteristics of the depolarizing influence of the field, as illustrated in
Extracellular stimulation typically generates a complex electric field in the tissue medium that is applied to the underlying neural processes as a distribution of extracellular potentials. As derived from the cable equation, the second derivative of the extracellular potentials along each process will typically produce both transmembrane and axial currents that will be distributed throughout the neuron. In turn, each neuron exposed to the applied field will typically experience both inward and outward transmembrane currents and regions of depolarization and hyperpolarization. These theoretical predictions have been verified in numerous experimental preparations demonstrating the differences between anodic, cathodic, and bipolar stimulation on the ability to both activate and block neural activity with extracellular stimulation.
Analysis of the effects of DBS is complicated by our limited understanding of the response of neurons near the electrode to the applied fields. Addressing the effects of high frequency DBS presents investigators with a paradox of how stimulation (traditionally thought to activate neurons) can result in similar therapeutic outcomes as lesioning target structures in the thalamus or basal ganglia. There exist two general philosophies on the effects of DBS: 1) DBS is believed to generate a functional ablation by suppressing or inhibiting the structure being stimulated or 2) DBS is believed to result in activation patterns in the stimulated network that override pathological network activity. Our model results support the latter theory by showing with detailed models and therapeutically effective stimulation parameters that axonal elements are activated over a large volume of tissue surrounding the electrode.
Experimental investigation on the effects of STN DBS has implicated activation of large diameter fiber tracks with therapeutic stimulation parameters. Predictions of the volume of tissue affected by DBS, using current-density calculations, have suggested that axonal elements would be activated over a 2.5 mm radius of the electrode contact using a −3 V stimulus. However, current-density is not directly related to the neural response to stimulation, and typically has a non-uniform distribution on DBS electrode contacts. A scaled version of the derivative of the current-density, Δ2Ve, represents a value that more accurately quantifies the stimulating influence of the electric field. Using Δ2Ve in combination with tissue electrical properties derived from DTI we predict that −3V STN DBS can activate axonal elements in STN, ZI, H2, and IC spreading as far as 4 mm from the electrode contact, as illustrated in
2. Examples of Using a Model
At 302, in one example, diffusion tensor imaging (DTI) data is obtained (this may occur at 300, such as where a DTI MR imaging modality is used at 300). In one example, the DTI data is obtained from the same patient being analyzed. Alternatively, “atlas” DTI data is obtained from at least one other patient. If atlas DTI data from another patient is used, it is typically spatially scaled to correspond to the anatomic size and shape of the patient being analyzed. In one example, the atlas DTI data is based on a composite from more than one other patient. The composite atlas DTI data typically spatially scales DTI data from the different patients before combining into the composite DTI atlas. The atlas DTI data avoids the need to obtain DTI data from the particular patient being analyzed. This is useful, for example, when a non-DTI imaging modality is used at 300.
At 304, a tissue conductivity model is created for all or part of the anatomic volume. The tissue conductivity model is typically a non-uniform spatial distribution. Such a model more accurately represents inhomogeneous and anisotropic characteristics of the tissue anatomy. For example, the conductivity of brain tissue varies from one brain region to another. Moreover, conductivity of the nervous system is preferential to a particular direction that is also dependent on the particular location in the brain. In one example, a non-uniform tissue conductivity model is created by transforming the DTI data into conductivity data, such as by using the linear transform techniques discussed above with respect to Equation 2.
It should be noted that it is not required to obtain non-uniform tissue conductivity data using DTI. There exist several alternatives to using DTI based approximations for the anisotropic and inhomogeneous tissue properties for the patient specific finite element volume conductor model. One example technique would be a simple designation of a white matter and a grey matter conductivity tensor. These two universal conductivity tensors could then be applied to the nodes of the FEM mesh using co-registration with the anatomical MRI. In this manner, the individual voxels of the MRI data are designated as either white matter or grey matter using post-processing image analysis. Then, each such voxel is assigned a conductivity dependent on whether it was classified as white matter or grey matter, which white matter voxels having a different conductivity value than grey matter voxels. A second example technique would define individual conductivity tensors for designated brain regions (e.g., nuclei, sub-nuclei, fiber tracts, etc.). This method would allow for a more detailed representation of the tissue electrical properties than the first example technique. The conductivity tensor of each designated brain region is defined, in one example, using explicit experimental tissue impedance results and anatomical information provided by a human brain atlas. In this technique, the anatomical MRI is sub-divided into different designated brain regions on a voxel-by-voxel basis using post-processing image analysis. The appropriate conductivity tensors for each designated brain region is then co-registered with the nodes of the FEM mesh.
At 306, a finite element model (FEM) is created using the conductivity data obtained at 304. In one example, the FEM model uses a default boundary condition that is appropriate for a typical electrode contact morphology. However, in another example, the FEM model includes an electrode-specific boundary condition that is tailored to the morphology of a particular electrode contact or contacts to be used in the DBS or other procedure. The FEM model provides for non-uniform conductivity in the tissue, such as by using a DTI-derived other conductivity value at each node in the FEM mesh. The FEM model may include aspects that are not obtained from the DTI-derived data. In one such example, the FEM mesh models a thin encapsulation sheath about the electrode lead body, as discussed above, which is not derived from the DTI data.
At 308, in one example, the FEM is solved for the electric potential distribution or the second difference (Δ2V) of the electric potential distribution, as discussed above, such as by using FEM solver software. In one example, the FEM is solved for a normalized stimulation amplitude of 1V. In another example, for a different electric stimulation amplitude, the resulting electric potential distribution (or second difference of the electric potential distribution) is multiplied by a scale ratio of the different electric stimulation amplitude to the normalized electric stimulation amplitude.
At 310, a volume of activation (VOA) or other volume of influence is calculated, in one example, using the second difference of the electric potential distribution. The VOA represents the region in which any neurons therein are expected to typically be activated, that is, they are expected to generate propagating action potentials at the stimulus frequency in response to the electrical stimulation delivered at the stimulation electrode contact. Conversely, neurons outside the VOA are expected to typically remain unactivated in response to the electrical stimulation. In one example, a particular threshold value of the second difference of the electric potential distribution defines the boundary surface of the VOA.
As discussed above, the particular threshold value defining the boundary of the VOA is determined as follows. First, model neuronal elements are positioned relative to the electrode using known neuroanatomical information about specific fiber pathways and nuclei of interest near the electrode. These generalized positions of the model neuronal elements are then refined, such as by using explicit “patient-specific” information provided in the DTI or anatomical MR imaging data. For example, the DTI imaging data describes the inhomogeneous and anisotropic tissue properties near the electrode. In this example, such DTI imaging data is used to explicitly define one or more axonal trajectories, if needed, or to help define nuclear boundaries specified in the anatomical MR.
A model of these neurons is then created. In one example, the neurons are modeled using an axon model, which is a simplified form of a neuron model. An example of an axon model is described in Cameron C. McIntyre et al., “Modeling the Excitability of Mammalian Nerve Fibers: Influence of Afterpotentials on the Recovery Cycle,” J. Neurophysiology, Vol. 87, February 2002, pp. 995-1006, which is incorporated by reference herein in its entirety, including its disclosure of axon models. In another example, a more generalized neuronal model is used, an example of which is described in Cameron C. McIntyre et al., “Cellular Effects of Deep Brain Stimulation: Model-Based Analysis of Activation and Inhibition,” J. Neurophysiology, Vol. 91, April 2004, pp. 1457-1469, which is incorporated by reference herein in its entirety, including its disclosure of neuronal models. The neuron model describes how the neurons will respond to an applied electric field, that is, whether the neuron will fire and whether the neurons will generate a propagating action potential.
In one example, using this neuron model to simulate how the neurons (located as determined from the DTI-derived conductivity data, in one example) behave, the threshold value of the second difference of electric field that will result in such propagating action potentials is calculated. The stimulating influence of the electric field is applied to the model neurons to define a threshold value. This threshold value is then used to define the boundary of the VOA in the non-uniform conductivity tissue, as discussed above.
It should be noted that the neuron model may depend on one or more of the electrical parameters of the DBS stimulation being modeled. For example, the stimulation pulsewidth will affect the neuron response. Therefore, in one example, the neuron model is tailored to a specific value for one or more DBS stimulation parameters.
It should also be noted that calculation of explicit threshold criteria for each patient is not required. For example, in a more generalized situation, threshold criteria will have already been determined using the detailed neuron models under a wide variety of different stimulation conditions. Once these threshold criteria have been determined, they need not be re-determined for each subsequent patient.
It should also be noted that using a threshold criteria upon the second difference of the potential distribution in the tissue medium is a simplified technique for quickly determining a VOA or other volume of influence. The intermediate step of using the second difference of the potential distribution is not required. In an alternate example, the FEM model of is directly coupled to a detailed neuron model, such as a multi-compartment neuron model that is oriented and positioned in the FEM model to represent at least one actual nerve pathway in the anatomic volume.
At 312, the calculated VOA region is displayed, such as on a computer monitor. In one example, the VOA is displayed superimposed on the displayed imaging data or a volumetric representation derived from such imaging data. In another example, an anatomic boundary or other representation of an anatomic structure is superimposed on the VOA and imaging data or the like. The anatomic boundary data is typically obtained from an atlas of brain anatomy data, which can be scaled for the particular patient, as discussed above. Alternatively, the anatomic representation is extracted from the imaging data for the patient being analyzed. In one example, the anatomic representation is a line depicting one or more boundaries between particular nucleus structures or other regions of the brain, such as the STN, IC, or ZI illustrated above in
In any case, by viewing a representation emphasizing one or more brain regions displayed together with the VOA, the user can then determine whether a particular anatomic region falls within or outside of the modeled VOA. The user may want a particular anatomic region to be affected by the DBS, in which case that region should fall within the modeled VOA. Alternatively, the user may want a particular region to be unaffected by the DBS, such as to avoid certain unwanted DBS stimulation side effects, as discussed above. This evaluation of whether the VOA is properly located can alternatively be performed by, or assisted by, a computer algorithm.
For example, the computer algorithm can evaluate various VOAs against either or both of the following input criteria: (a) one or more regions in which activation is desired; or (b) one or more regions in which activation should be avoided. In one example, at 314, the computer algorithm creates a score of how such candidate VOAs map against desired and undesired regions. In one example, the score is computed by counting how many VOA voxels map to the one or more regions in which activation is desired, then counting how many VOA voxels map to the one or more regions in which activation is undesired, and subtracting the second quantity from the first to yield the score. In another example, these two quantities may be weighted differently such as, for example, when avoiding activation of certain regions is more important than obtaining activation of other regions (or vice-versa). In yet another example, these two quantities may be used as separate scores.
At 316, the score can be displayed to the user to help the user select a particular VOA (represented by a particular electrode location and parameter settings). Alternatively, the algorithm can also automatically select the target electrode location and parameter settings that provide the best score for the given input criteria.
In one example, the VOA is displayed on a computer display monitor of an image-guided surgical (IGS) workstation, such as the StealthStation® from the Surgical Navigation Technologies, Inc. (SNT) subsidiary of Medtronic, Inc., for example. The VOA can be displayed on the IGS workstation monitor with at least one of the imaging data representing the anatomic volume, the target electrode location, a burr hole or other anatomic entry point, a trajectory between the anatomic entry point and the target electrode location, or an actual electrode location.
In one IGS workstation example, the displayed VOA corresponds to a target electrode location. Another IGS workstation example provides an intraoperatively displayed VOA corresponds to an actual electrode location of an electrode being introduced along the trajectory. The VOA is recomputed and redisplayed as the electrode is being introduced along the trajectory, such as by using position information tracking the position of the electrode being introduced. In one example, various VOAs along the trajectory are pre-computed, and the particular VOA is selected for display using the tracked position of the electrode as it is being introduced.
After the electrode is positioned at the target location, it is typically secured in place, such as by using a lead immobilizer located at the burr hole or other anatomic entry point. There remains the challenging task of adjusting the DBS stimulation parameters (e.g., the particular electrode contact(s) of a plurality of electrode contacts disposed on the same DBS leadwire, pulse amplitude, pulsewidth, electrode “polarity” (i.e., monopolar or bipolar electrode return path), electrode pulse polarity (i.e., positive or negative), frequency, etc.). In one example, the IGS workstation or a DBS pulse generator programmer includes the above-described VOA methods to assist the user in selecting an appropriate combination of DBS stimulation parameters, such as by using the scoring techniques discussed above.
3. Application in a Patient-Specific Neural Stimulation Modeling System
One application of the above-described neural response modeling techniques is in a patient-specific neural stimulation modeling system (PSNSMS), which can be implemented as a software package that, in one example, can be integrated into an IGS workstation or any other desired computer implementation. The PSNSMS allows interactive manipulation of patient-specific electrical models of the brain for analysis of brain stimulation methods. This provides a virtual laboratory for surgeons, technicians, or engineers to optimize or otherwise adjust neural stimulation treatment, such as by varying electrode position, stimulation protocol, or electrode design. In one example, the PSNSMS integrates data processing, numerical solution and visualization into one cohesive platform. In one example, the PSNSMS uses a modular framework that incorporates anatomical or functional magnetic resonance images, 3D geometric models of individual brain nuclei, volume conductor models of the electric field generated by the stimulation, biophysical models of the neural response to the stimulation, numerical solutions of the coupled electric field and neuron models, and 3D visualization of the model results and imaging data. Among other things, the PSNSMS outputs a volume of influence (neural activation or neural inhibition) generated by the stimulating electrode for a given position in the brain and given stimulation parameters.
Benefits of the PSNSMS may include, among other things: (1) pre-operative targeting of an optimal or desirable neural stimulation electrode position or trajectory in the brain tissue medium; (2) intra-operative monitoring or visualization of electrode position or trajectory and stimulation effects as a function of the stimulation parameters; (3) post-operative adjustment or optimization of one or more stimulation parameters for therapeutic benefit given knowledge of the actual electrode position in the brain; or (4) a design tool for evaluating or testing different electrode designs, such as for a given anatomical target.
Existing techniques for pre-operatively targeting specific nuclei for neurostimulation using magnetic resonance imaging data only account for certain anatomical considerations. They typically ignore the electric field generated by the stimulation and the subsequent neural response to the stimulation. Existing techniques for intra-operatively monitoring the electrode position in the brain, based on the spontaneous electrical activity of neurons surrounding the electrode, require highly skilled neurophysiologists to interpret the data. Moreover, such techniques are not linked with 3D visualization of the surrounding neuroanatomy. Furthermore, they do not enable prediction of the effects of stimulation as a function of the stimulation parameters. Existing techniques for defining effective stimulation parameter values typically rely on trial and error. They typically do not explicitly take into account the anatomical position of the electrode or the neural response to stimulation as it depends on changes in the stimulation parameters. Moreover, they typically do not use any optimization strategies to define the stimulator parameter settings.
The PSNSMS addresses these and other limitations. In one example, the PSNSMS uses a finite element model (FEM) of the electric field generated by the stimulation. In one example, the tissue electrical properties of the FEM are based on diffusion tensor magnetic resonance imaging analysis, also referred to as diffusion tensor imaging (DTI). DTI permits explicit characterization of the inhomogeneous and anisotropic tissue properties near a given electrode position. The inhomogeneous and anisotropic tissue properties distort the electric field. Therefore, they are important to consider when addressing the neural response to the stimulation.
In one example, the electric field model is then coupled to electrical models of individual neurons to predict their response to the applied stimulation and determine a volume of tissue that is directly influenced by the stimulation. In another example, simplifying assumptions allow the volume of activation (VOA) to be obtained directly from the electric field model using the second difference of the potential distribution in the tissue medium, as discussed above.
The PSNSMS also allows integration of MR imaging data, 3D anatomical volumes, neural stimulation electrode trajectory, and 3D neural stimulation response volume in a single platform or package. This platform or package can be used for, among other things, pre-operative targeting, intra-operative monitoring, post-operative stimulation parameter adjustment or optimization, or electrode design. One example of such a package is a image-guided surgical (IGS) workstation, which typically displays voxel data obtained from MR or CT images, and to which a display of a modeled 3D neural stimulation response volume of influence (or other information obtained from a modeled 3D neural stimulation response volume of influence) has been added.
B. Exemplary Methods
In one example, the PSNSMS allows, among other things, capture of the detailed interaction between the electric field and underlying non-uniform tissue medium. This enables more accurate estimation of the spatial extent of neural activation generated by one or more electrodes implanted in the nervous system. In one embodiment, the PSNSMS includes the following components: (1) a volume conductor electric field model such as a FEM mesh, which includes a model of the stimulating electrode and of any inhomogeneous and anisotropic electrical properties of nearby tissue; (2) one or more multi-compartment electrical models of individual neurons whose positions can be specified within the electric field (e.g., using anatomically-derived conductivity data to ascertain the locations of neural pathways) and their response to stimulation can be quantified; (3) integration of functional or anatomical imaging data into a visualization platform that can be combined with the electric field modeling results; or, (4) techniques to determine a desired or optimal electrode position or one or more desired or optimal stimulation parameters on a patient-specific basis.
In the example of
The example of
Our example demonstration of PSNSMS is based on deep brain stimulation (DBS) of the subthalamic nucleus (STN), but the concepts described in this document are transferable to any electrode design or to stimulation of any region of the nervous system. In one example, one or more portions of the PSNSMS is constructed using the shareware package SCIRun with BioPSE (Scientific Computing and Imaging Institute, University of Utah), which provides an integrated environment for data manipulation, analysis, and interactive visualization.
C. Volume Conductor Electric Field Model Example
In one example, detailed patient-specific electric field models of central nervous system stimulation were developed using anatomical and diffusion tensor magnetic resonance data (DTI).
In one example, the DTI data was used to estimate the inhomogeneous and anisotropic tissue conductivity properties on a patient-specific basis and this information was integrated into the FEM. As described above, the electrical conductivity tensor (σ) was determined from the diffusion tensor (D) at each voxel of the DTI, such as by using Equation 2.
After a pre-operative electrode target location or post-operative implanted electrode location is determined, in one example, the volumetric conductivity data from the DTI (also referred to as a DTI voxel map) is co-registered with the FEM illustrated in
σ′=RσRT (Eq. 3)
where R is the rotation matrix for the image transformation defined by α and β.
After this transformation, each node of the FEM mesh is assigned a conductivity tensor that is mapped to its corresponding location within the DTI voxel map. In one example, the FEM mesh illustrated in
After the FEM is defined with the appropriate tissue conductivity data, appropriate boundary conditions are set, as discussed above. Then, Equation 1 is solved to determine the electric potential distribution generated in the tissue medium. Equation 1 is typically solved using one of two solvers, depending on the characteristics of the stimulation waveform used, an example of which is illustrated in voltage amplitude vs. time graph of
D. Example of Quantifying the Neural Response to Stimulation
Knowing the potential distribution in the tissue medium alone is not enough to predict the neural response to stimulation. Therefore, in one example, we use one or more multi-compartment cable models of individual neurons to address the neural response to the stimulation. Such neuron models represent electrically equivalent circuit representations of physiological neural signaling mechanisms. The models typically include an explicit representation of the complex neural geometry and individual ion channels that regulate generating of action potentials. The neuron model geometries are typically broken up into many (e.g., hundreds) of compartments and are co-registered within the FEM mesh. This allows calculation of the extracellular potentials from the applied electric field along the complex neural geometry. After the extracellular potentials are determined for each neural compartment as a function of time during the applied stimulation, for each neural position relative to the electrode, the model neuron is used to test whether the applied stimulus exceeded the neural threshold that triggers an action potential. The neural response to extracellular stimulation is dependent on several factors, such as, for example: (1) the electrode geometry; (2) the shape of the electric field (as determined by the inhomogeneous and anisotropic bulk tissue properties); (3) the neuron geometry; (4) the neuron position relative to the stimulating electrode; (5) the neuron membrane dynamics; and, (6) the applied stimulation parameters (e.g., stimulus waveform, stimulation frequency, etc.).
In one illustrative example, we used the 5.7 μm diameter double cable myelinated axon model described in Cameron C. McIntyre et al., “Modeling the Excitability of Mammalian Nerve Fibers: Influence of Afterpotentials on the Recovery Cycle,” J. Neurophysiology, Vol. 87, February 2002, pp. 995-1006, which is incorporated herein by reference in its entirety. (Alternatively, instead of using an axon model, a more detailed neuronal model could be used, such as described in Cameron C. McIntyre et al., “Cellular Effects of Deep Brain Stimulation: Model-Based Analysis of Activation and Inhibition,” J. Neurophysiology 91: 1457-1469 (2004), which is incorporated by reference herein in its entirety). We incorporated this model into our STN DBS FEM to quantify the neural response to stimulation. By positioning the axon in different locations relative to the electrode and modulating the stimulation parameters one can determine the threshold stimulus necessary to activate the neuron. Likewise, for a given stimulation parameter setting (pulse duration, amplitude, frequency), the threshold characteristics of the model neuron can be determined as a function of space around the electrode. This information defines of a volume of tissue for which the neural activation threshold is exceeded for the particular stimulation parameter setting. This volume of tissue is referred to as the volume of activation (VOA). In one example, a further simplification is made by determining a threshold value of the second difference of the potential distribution, which is representative of neural activation for a given stimulation parameter setting, as discussed above and as illustrated in
When using PSNSMS to pre-operatively characterize stimulation effects, assumptions are typically made as to the appropriate model parameter values used to predict the volume of activation. However, during post-operative use, the PSNSMS model can be fit to patient-specific experimental threshold results. The tissue conductivity and electrode localization for each patient-specific FEM can be adjusted to match the clinically determined threshold stimulation results for activating major fiber tracts near the electrode. Detecting fiber tract activation may involve monitoring behavioral responses that are known to arise from such activation of specific fiber tracts. The clinical threshold to elicit these behavioral responses is determined. These fiber tracts can be explicitly visualized on the DTI voxel map. The location and trajectory of particular fiber tracts can be directly integrated into PSNSMS by positioning the axon models along the appropriate anatomical trajectory in the FEM. Three general variables can be adjusted to fit the FEM to the experimental data. First, the conductivity of the encapsulation layer about the electrode can be adjusted (e.g., 0.2 S/m<σencap<0.1 S/m) to fit the FEM to the experimental data. Alterations in this variable modulate the electrode input impedance. Such adjustments can be guided by clinical data from the stimulator programming unit. Second, the ratio of effective extracellular conductivity and the effective extracellular diffusivity (0.6<σe/de<0.8 (S−s)/mm2) can be adjusted. Altering this variable scales the absolute value of the conductivity tensor and modulates the stimulus amplitudes needed for axonal activation. A third variable is the X, Y, Z position of the electrode relative to the tissue medium. We expect about 1 mm error in our MR-based electrode localization due to the metallic distortion artifact generated by the electrode in the MR image. Therefore, in one example, we allow the electrode to be shifted by a maximum of 0.5 mm in any direction to allow convergence between the model-predicted threshold data and the clinical threshold data.
E. Example of Integrating Stimulation Modeling Results and Anatomic Imaging Data
Calculating the volume of activation as a function of the electrode location and stimulation parameters represents one component of PSNSMS. This provides even greater utility when it is integrated with patient-specific anatomical data. Anatomical data derived from MRI is commonly used to target stereotactic neurosurgical procedures. In one example, however, the PSNSMS integrates and displays the anatomical data with volume of activation data, as illustrated by the computer display screenshot of
In one example, the PSNSMS includes a patient-specific brain atlas. Such atlases can be generated from the pre-operative anatomical MR images using techniques originally described by Gary E. Christensen et al., “Volumetric Transformation of Brain Anatomy,” IEEE Trans. on Medical Imaging, Vol. 16, No. 6, pp. 864-877, December 1997, which is incorporated herein by reference in its entirety. However, any variety of morphing algorithms could be used. One suitable algorithm includes a nonlinear transformation to register one MRI (the patient-specific image) to a second pre-labeled target MRI that serves as a canonical atlas for particular regions of the brain. Segmentation of the patient-specific MRI is achieved by using the inverse of this transformation to warp the canonical atlas back onto the patient's 3D MR image. In one example, the registration procedure operates in two stages. In the first stage, a low-dimensional registration is accomplished by constraining the transformation to be in a low-dimensional basis. In one example, the basis is defined by the Green's function of the elasticity operator placed at pre-defined locations in the anatomy and the eigenfunctions of the elasticity operator. In the second stage, high-dimensional large transformations are vector fields generated via the mismatch between the template and target-image volumes constrained to be the solution of a Navier-Stokes fluid model. The result of these transformations is a 3D brain atlas matched to the individual patient with specific volumes representing pre-labeled target nuclei. The 3D surface data derived from the patient-specific brain atlas is then co-registered and, in one example, is displayed with the electrode and volume of activation data, as illustrated in the example of
F. Example of Model-Based Selection of Patient-Specific Target Electrode Locations or Stimulation Parameter Settings
One purpose of the PSNSMS is to determine optimal or desirable preoperative electrode locations or post-operative optimal or desirable stimulation parameters settings on a patient-specific basis. This typically involves determining a target volume of tissue that should be activated by the stimulation. In the PSNSMS, the geometry of this target VOA is typically determined based on the patient-specific 3D brain atlas. For example, in the case of STN DBS for Parkinson's disease, current anatomical and physiological knowledge indicate that the target volume of tissue is the dorsal half of the STN. Therefore, in this example, for each patient-specific 3D brain atlas we determine a target VOA defined by the dorsal half of the STN. We then determine test VOAs generated by a range of electrode positions within the STN and/or a range of stimulation parameter settings for each of those electrode locations. These test VOAs are then compared to the target VOA. The electrode position and/or stimulation parameter setting that generates a test VOA that most closely matches the target VOA is provided as the model-selected electrode position and/or stimulation parameter setting.
In one variant of this selection process, engineering optimization is used to assist the selection process. Examples of possible constraints on the selection process include one or more of minimizing charge injection into the tissue, minimizing spread of the test VOA outside of the target VOA, maximizing overlap between the test VOA and target VOA, limiting the stimulus amplitude to being greater than −10 V and less then 10V, limiting the stimulus pulse duration to being greater than 0 and less than 450 ms, limiting the stimulation frequency to being greater than 0 and less than 185 Hz. In one such example, limits on the stimulation parameters are determined by the output of the current clinical stimulator. Therefore, if new technology provides new output limits, our model limits could be refined to reflect these changes. In a further example, the engineering optimization uses one or more penalty functions, which can be applied for test VOAs that spread into neighboring anatomical structures that are known to induce side effects.
When using PSNSMS pre-operatively, in one example, both the electrode location and stimulation parameters can be varied to determine test VOAs that match the target VOA. This helps determine a pre-operative target for stereotactic neurosurgical implantation of the electrode.
When using PSNSMS post-operatively, in one example, the modeled electrode position in the tissue medium is established using the actual implanted electrode location. Then, one or more stimulation parameters are varied to determine test VOAs, which are compared to the target VOA to determine which test VOA (and hence, which parameter setting(s)) obtain the closest match to the target VOA. This indicates which chronic stimulation parameter settings maximize or otherwise provide the desired therapeutic benefit.
It is to be understood that the above description is intended to be illustrative, and not restrictive. For example, the above-described embodiments (and/or aspects thereof) may be used in combination with each other. Many other embodiments will be apparent to those of skill in the art upon reviewing the above description, and aspects of described methods will be computer-implementable as instructions on a machine-accessible medium. The scope of the invention should, therefore, be determined with reference to the appended claims, along with the full scope of equivalents to which such claims are entitled. In the appended claims, the terms “including” and “in which” are used as the plain-English equivalents of the respective terms “comprising” and “wherein.” Also, in the following claims, the terms “including” and “comprising” are open-ended, that is, a system, device, article, or process that includes elements in addition to those listed after such a term in a claim are still deemed to fall within the scope of that claim. Moreover, in the following claims, the terms “first,” “second,” and “third,” etc. are used merely as labels, and are not intended to impose numerical requirements on their objects.
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|U.S. Classification||600/407, 703/2, 607/116, 703/13|
|Cooperative Classification||A61B2090/364, A61B90/37, G06F19/3437, G06F19/321, A61N1/0534, A61N1/0539, A61N1/36146, A61N1/36082, A61N1/36135|
|European Classification||G06F19/34H, A61N1/36Z, A61N1/36Z3E|
|Oct 2, 2014||FPAY||Fee payment|
Year of fee payment: 4
|Oct 2, 2014||SULP||Surcharge for late payment|